Porous metallic scaffold for tissue ingrowth

ABSTRACT

The invention relates to implantable medical devices, particularly, to porous structures for such devices. In one aspect, the invention provides a porous metal scaffold comprising a porous metal network having pores defined by metal webs, the metal webs covered with at least one layer of metal particles bonded to the metal webs. In other aspects, the invention provides methods of forming porous scaffolds. In one such aspect, the method includes providing a polymer foam; forming a skin of biocompatible metal on the polymer foam by low temperature arc vapor deposition; and heating the polymer foam and the metal skin above the decomposition temperature of the polymer foam in an inert gas atmosphere; thereby the polymer foam decomposes producing a green metal foam. In yet other aspects, the invention provides methods of improving stability of porous scaffolds.

FIELD OF THE INVENTION

The present invention relates to implantable medical devices, and inparticular to porous scaffolds for implantable medical devices,especially orthopedic implants, methods of forming such scaffolds, andmethods of producing such devices.

BACKGROUND OF THE INVENTION

The use of orthopedic implants is the result of deterioration of humanbone structure, usually because of various degenerative diseases, suchas osteoarthritis. In recent years, a variety of implantable orthopedicdevices had been developed. Typically, the failed bone structure isreplaced with an orthopedic implant that mimics the structure of thenatural bone and performs its functions.

Orthopedic implants are constructed from materials that are stable inbiological environments and withstand physical stress with minimaldeformation. Such materials must possess strength, resistance tocorrosion, biocompatibility, and good wear properties. Also, theimplants include various interacting parts, which undergo repeatedlong-term physical stress inside the body.

A breakdown of a permanently installed implant leads to pain, limitationon the range of motion, and may require a replacement of the implant.For these reasons, among others, the bone/implant interface and theconnection between various parts of the implant must be resistant tobreakdown. It is especially important since installation of anorthopedic implant often involves extensive and difficult medicalprocedure, and therefore replacement of the installed implant is highlyundesirable.

The requirements for the useful life of the implant continue to growwith the increase in the life expectancy. The strength and longevity ofimplants in large part depend on the bone/implant interface. Variousmethods of connection are known in the art. For example, a hip joint isa ball-in-socket joint, and includes a rounded femoral head and acup-like socket (acetabular cup) located in the pelvis. The surfaces ofthe rounded femoral head and the acetabular cup continually abrade eachother as a person walks. The abrasion creates stress on the bones thatbear the acetabular cup and the femoral head. If the femoral head or theacetabular cup is replaced with an implant, this stress must be welltolerated by the implant's bearing surfaces to prevent implant failure.

FIG. 1 shows a typical hip replacement system that includes anacetabular cup prosthetic assembly 10 and a femoral prosthesis 20.Generally, the acetabular cup implant 10 includes a bone interface shell11 and a socket bearing insert 12. The femoral prosthesis 20 includes afemoral stem 21 and a femoral head in the form of a ball 22, which movesinside the socket insert 12 of the acetabular cup implant 10. Thefemoral ball 22 usually has a polished surface to maintain a LOWfriction interface with the surface of the socket insert 12 of theacetabular cup 10. The stem section 21 is inserted into the interior ofthe femur and may have a bone interface surface 26.

The socket insert 12 is usually made from a plastic material such aspolyethylene or ultra high molecular weight polyethylene (UHMWPE), butmay be of any biocompatible material that has sufficient strength andwear resistance to withstand the pressures and abrasive nature of thejoint. The socket insert 12 is typically held in the shell 11 by aseries of locking grooves or notches. In turn, the complete acetabularcup implant 10 may be attached to the patient's pelvis by a series oflocking grooves, pins or screws 29. Alternatively, the acetabular cupimplant 10 may be press-fit by being driven into the patient'sacetabulum with a proper impaction tool without the fixing pins insituations where patient-related criteria are met. This method avoidsthe use of bone cement. The shell 11 is typically made from a metal suchas titanium or cobalt-chrome alloy, and has a bone interface surface 16.

In use, the bone interface surfaces 16 and 26 must bear a significantlateral and axial stress. The increased requirements for useful life ofthe implant make it especially important that these surfaces toleratesuch stress. The prior art takes several approaches to this problem.

Thus, the entire acetabular cup implant 10, including both the shell 11and the socket insert 12, may be cemented to the acetabulum or the cupmay be produced as a single piece from ultra high molecular weightpolyethylene and anchored into the acetabulum with bone cement. Anotherway to improve the longevity of orthopedic implants is to provide aporous bone interface surface to receive ingrowth of bone tissue therebybinding the natural bone to the implant. The bone ingrowth into thevoids of the porous bone interface layer provides skeletal fixation forthe implants used for replacement of bone segments. In addition tolateral and axial strength enhancement, the bone ingrowth improvesbiocompatibility of the implant and is even believed by some to promotepositive biochemical changes in the diseased bone. To implement thisapproach, it is important to develop methods of constructing porousouter layers on the bone interface surfaces of implants.

Orthopedic implants with porous bone interface surfaces have beenstudied extensively over the last twenty years. It has long been knownthat the success in facilitating the ingrowth is related to the porecharacteristics of the bone interface surfaces, such as pore size andpore volume. For example, it is known that the bone ingrowth may bealmost entirely non-existent if the porous layer has pore sizes of lessthan 10 μm, and that pore sizes greater that 100 μm facilitate theingrowth.

In view of the strength and longevity requirements, the implants aretypically made of biocompatible metals, such as titanium orcobalt-chrome alloy. Thus, one of the challenges is to provide metallicorthopedic implants having porous metallic bone interfaces with highporosity. Another challenge is to provide an integrated bond between theporous layer and the underlying solid substrate, such as the surface 16and the bulk of the shell 11, respectively, of the acetabular cupimplant 10 shown in FIG. 1.

Certain orthopedic implants having porous bone interface surfaces, andrelated methods of making such implants have been patented. U.S. Pat.No. 5,282,861 describes an open cell tantalum structures for boneimplants having pore volume of from 70 to 80%. The open cell tantalumstructures of the '861 patent are formed by chemical vapor deposition oftantalum on a carbon skeleton. The resulting structures have a carboncore and a tantalum outer surface.

U.S. Pat. No. 6,087,553 describes tantalum/polyethylene compositessuitable for use in orthopedic implants. The composites have a porevolume of 50 to 90%. The implants produced from the composites of the'553 patent are not modular and not metal-backed.

In general, methods of producing high pore volume metals are known inthe art. U.S. Pat. No. 5,976,454 describes a process for producingnickel foam for use in making battery electrodes. The porosity of thefoam is over 90%, but it is produced by a method that is in manyrespects not suitable for producing foams of biocompatible metalstypically used in making implants, such as tantalum or titanium.

U.S. Pat. No. 5,926,685 describes a method of forming an implant havinga porous outer surface by using an organic binder compound to enhancethe binding between the porous surface layer and the implant. The binderand metal particles that would form the porous layer are mixed and themixture is placed in contact with a solid surface of the metallicimplant. Then, the particles (pre-cursor of the porous layer) are boundto each other and to the solid surface of the implant via a sinteringprocess. The '685 patent does not describe production of a metal foam asa pre-cursor to the porous layer. Also, the '685 patent does notdescribe the porosity of the porous layer.

Therefore, there exists a continuing need for implantable medicaldevices, especially orthopedic implants, having porous surfaces, blocks,layers or other porous structures for interfacing with bones and/orother tissue, with the porous structures having a variety of desirablecharacteristics, including high porosity, uniform pore size, and highstrength.

SUMMARY OF THE INVENTION

Various aspects of the present invention address this need. Thus, inaccordance with one aspect, the invention provides a porous metalscaffold for use in an implantable medical device comprising a porousmetal network having pores defined by metal webs, the metal webs coveredwith at least one layer of metal particles bonded to the metal webs.Preferably, the metal webs of the porous metal scaffold may form acontinuous inner skeleton. The pore size of the porous scaffold may bevaried by bonding additional layers of metal particles to the at leastone layer of particles. Also, changing a size of the metal particles mayalso vary the pore size of the porous scaffold.

Preferably, the bonding between the metal webs and the metal particlesis accomplished by sintering the metal particles to the webs. Also,preferably, the metal webs have partially hollow cores. The hollow coresof the metal webs may be surrounded by an outer web wall that hasopenings therein.

The pore size of the porous scaffold may range from 100 μm to 1000 μm.The pore volume may range from 50% to 90%. The scaffold may be formedinto a shape having a thickness of 0.5 mm to 5 mm.

Preferably, the porous metal scaffold is bonded to a solid metalsubstrate. Also, preferably, the porous metal scaffold is directlybonded to the solid metal substrate. The metal scaffold may be sinteredto the solid metal substrate. The scaffold may include a plurality ofpores having a size greater than about 100 μm. The metal particles mayhave a size from 40 μm to about 80 μm. The metal of the particles ispreferably selected from the group consisting of titanium, titaniumalloy, cobalt chrome alloy, niobium and tantalum. The web metal is alsopreferably selected consisting of titanium, titanium alloy, cobaltchrome alloy, niobium and tantalum. The metal substrate may be part ofan orthopedic implant.

In accordance with another aspect, the invention provides a method offorming a porous scaffold for use in an implantable medical device, themethod including:

a) providing a polymer foam having a pre-determined thickness and a poresize ranging from about 500 μm to about 2000 μm;

b) forming a skin of biocompatible metal on the polymer foam by lowtemperature arc vapor deposition;

c) heating the polymer foam and the metal skin above the decompositiontemperature of the polymer foam in an inert gas atmosphere; thereby thepolymer foam decomposes producing a green metal foam.

Preferably, the method of this aspect of the invention further includesthickening the green metal foam by applying a solution of a binder ontothe green foam, applying a metal powder having a pre-determined particlesize, and sintering the foam, thus producing a final metal foam having apre-determined pore size. The thickening of the foam may be repeateduntil the final metal foam has the pre-determined pore size.

Preferably, the pre-determined thickness of the polymer foam is betweenabout 0.5 mm and about 10 mm, more preferably, between about 1 mm andabout 5 mm, yet more preferably, between about 1 mm and about 2 mm. Thepreferred polymer foam is polyurethane foam. Preferably, the polymerfoam has a pore size ranging from about 900 μm to about 1100 μm.

Preferably, the metal skin has thickness between about 1 μm and about 50μm. More preferably, the polymer foam has a first side and a secondside, and the thickness of the metal skin is about 35 μm on the firstside and about 10 μm on the second side.

Preferably, the binder solution is an aqueous solution of methylcellulose. Also, preferably, the pre-determined particle size of themetal particles used to thicken the metal web is between about 20 μm andabout 100 μm, more preferably, between about 40 μm and about 80 μm.

Preferably, the pre-determined pore size of the final metal foam isbetween about 100 μm and about 1000 μm, more preferably, between about300 μm and about 500 μm.

The invention also provides the green metal foam and the final metalfoam produced by the method(s) of this aspect of the invention, as wellas any intermediate metal foam. Preferably, the pre-determined pore sizeof the final metal foam produced by such method(s) is between about 100μm and about 1000 μm, more preferably, between about 300 μm and about500 μm, and/or a pore volume from about 50% to about 90%, morepreferably, from about 60% to about 80%. The preferred final metal foamis made of titanium or titanium alloy.

The final metal foam produced by the method(s) of this aspect of theinvention may be attached to a solid metal substrate. Such final metalfoam may be included in the implantable medical device. The preferredimplantable medical devices are orthopedic implants. One preferreddevice is an acetabular cup implant.

The biocompatible metal of the metal skin formed in the method of theinvention may be titanium, titanium alloy, cobalt chrome alloy, niobiumor tantalum. Also, the final metal foam and/or the solid substrate ofthe orthopedic implant also may be made of titanium, titanium alloy,cobalt chrome alloy, niobium or tantalum. Preferably, the final metalfoam and the substrate are produced from titanium or titanium alloy.

In another aspect, the invention provides a method of forming a porousscaffold for use in an implantable medical device, that includes:

a) providing a first metal foam of biocompatible metal;

b) spraying an atomized mist of a binder solution on the first metalfoam, wherein said mist has an average droplet size ranging from about20 μm to about 80 μm;

c) delivering a plurality of metal particles to the metal foam;

d) bonding the metal particles to the first metal foam; wherebyproducing a second metal foam having smaller pore size than the firstmetal foam.

The steps (b), (c) and (d) may be repeated if desired. Preferably, themist is produced by an ultrasonic source. The preferred average dropletsize ranges from about 30 μm to about 40 μm. The preferred bindersolution is an aqueous solution of methyl cellulose. Preferably, themetal of the first metal foam and/or of the metal particles is titanium,titanium alloy, cobalt chrome alloy, niobium or tantalum.

The invention also provides the second metal foam produced by themethod(s) of this aspect of the invention. The preferred pore size ofthe second metal foam ranges from about 100 μm to about 1000 μm, morepreferably, from about 300 μm to about 500 μm.

In yet another aspect, the invention provides a method of forming aporous scaffold for use in an implantable medical device, the methodincluding:

a) providing a polymer foam having a pre-determined thickness and afirst pore size;

b) forming a metal skin network of biocompatible metal on the polymerfoam by low temperature arc vapor deposition;

c) decomposing the polymer foam in an inert gas atmosphere therebyforming a green metal foam;

d) pre-sintering the green metal foam;

e) contacting the pre-sintered metal foam with metal particles in thepresence of a binder;

f) bonding the metal particles to the pre-sintered metal foam;

whereby obtaining the porous scaffold having pores of a second poresize.

Preferably, the pre-determined thickness of the polymer foam is fromabout 0.5 mm to about 2 mm. Also, preferably, the first pore size isfrom about 900 μm to about 1100 μm. The preferred second pore size isfrom about 300 μm to about 500 μm.

The preferred inert atmosphere is argon atmosphere. Preferably, themetal particles and the pre-sintered foam are bonded by sintering. Thepreferred metal of the scaffold is titanium, titanium alloy, cobaltchrome alloy, niobium or tantalum. The preferred metal of the metalparticles is also titanium, titanium alloy, cobalt chrome alloy, niobiumand tantalum.

In yet another aspect, the invention provides a method of improvingstability of a porous scaffold in an orthopedic implant, that includes

a) providing a pre-cursor for the orthopedic implant, the pre-cursorincluding a body and a spaced member attached to the body, the spacedmember including a wall member spaced from the body and a spacer elementconnecting the wall member to the body thereby the spacer element, thebody, and the wall member define a recess;

b) attaching the porous scaffold to the body, wherein the porousscaffold has a pre-determined pore size and includes a first portion anda second portion, which extends into the recess;

c) filling the recess, including the second portion of the porousscaffold, with metal particles having a particle size smaller than thepore size of the porous scaffold thereby the pores of the second portionof the porous scaffold are filled with the metal particles;

d) sintering the implant pre-cursor, the metal particles, and theattached porous scaffold including the filled second portion therebyconverting the spaced member, including the filled recess to asubstantially solid metal block, including converting the filled secondportion of the scaffold to a substantially solid portion of thesubstantially solid metal block; whereby the substantially solid portionat least partially supports the first portion of the porous scaffold.

The method of this aspect of the invention may further includesubjecting the pre-cursor with the filled spaced member to a vibrationaltreatment before sintering. The invention also provides an orthopedicimplant produced by the method of this aspect of the invention.

In yet another aspect, the invention also provides a method of improvingstability of a porous scaffold in an acetabular cup implant, the methodincluding

a) providing a blank acetabular cup shell, including a body having a topsurface, the blank shell including a rim, the rim having a ledge and awall spaced from the body of the blank acetabular cup shell thereby theledge, the body, and the wall define a circular annular recess;

b) attaching the porous scaffold to the top surface of body, wherein theporous scaffold has a pre-determined pore size and includes a firstportion and a second portion, the second portion of the scaffoldextending into the circular recess;

c) filling the recess, including the second portion of the porousscaffold, with metal particles having a particle size smaller than thepore size of the porous scaffold thereby the pore of the second portionof the porous scaffold are filled with the metal particles;

d) bonding the particles to the ledge, the wall, the second portion ofthe porous scaffold, and to each other, thereby converting the rim to asubstantially solid metal block, including converting the filled secondportion of the scaffold to a substantially solid portion of thesubstantially solid block;

whereby the substantially solid portion at least partially supports thefirst portion of the porous scaffold.

Preferably, in the method of this aspect of the invention the bonding iseffected through sintering.

The method of this aspect of the invention may further include machiningthe substantially solid block into a desired shape. The invention alsoprovides the acetabular cup implant that includes an acetabular cupshell produced according to the method(s) of this aspect of theinvention.

DESCRIPTION OF THE DRAWINGS

A more accurate appreciation of the subject matter of the presentinvention and the various advantages thereof can be realized byreference to the following detailed description, which makes referenceto the accompanying drawings in which:

FIG. 1 shows a typical hip joint implant system;

FIG. 2 illustrates one of the embodiments of a porous metal scaffold ofthe invention suitable for use in implantable medical devices;

FIG. 3 shows a general functional block diagram of a method forproducing porous metal scaffolds in accordance with one of theembodiments of the invention;

FIGS. 4A and 4B show an example of a polyurethane shell that matches theshape of an acetabular cup shell in accordance with one embodiment ofthe invention;

FIG. 4C illustrates a titanium-coated polyurethane shell shown in FIGS.4A and 4B;

FIG. 5 shows scanning electron microscope (SEM) photographs of thetitanium-coated polyurethane foam at 25× and 1000× magnifications;

FIG. 6 shows a schematic front cross-sectional view of a blank shell ofan acetabular cup implant in accordance with the preferred embodiment ofthe invention;

FIG. 7 shows a schematic front cross-sectional view of an assemblage ofthe blank shell of an acetabular cup implant and the titanium-coatedpolymer foam in accordance with the preferred embodiment of theinvention;

FIG. 8 shows a scheme of the furnace set-up for decomposing polyurethanefoam in accordance with the preferred embodiment of the invention;

FIG. 9 shows a schematic front cross-sectional view of the blank shellof an acetabular cup implant having a green titanium foam in accordancewith the preferred embodiment of the invention;

FIG. 10A shows scanning electron microscope photographs of a greentitanium foam at 25× and 200× magnifications;

FIG. 10B shows scanning electron microscope photographs of apre-sintered titanium foam at 25× and 1000× magnifications;

FIG. 11 is a functional block diagram of the preferred variant of theweb thickening process in accordance with the preferred embodiment ofthe invention;

FIG. 12 shows a schematic front cross-sectional view of the blank shellof an acetabular cup implant having a thickened titanium foam inaccordance with the preferred embodiment of the invention;

FIG. 13 shows an SEM photograph of the final titanium foam shown at 25×magnification;

FIGS. 14A-14B illustrate a blank shell of an acetabular cup implanthaving a rim in accordance with the preferred embodiment of theinvention;

FIG. 14C shows the rim of the acetabular cup implant in accordance withthe preferred embodiment of FIGS. 14A-14B;

FIG. 15A-15C illustrate an acetabular cup shell having a thickenedtitanium foam, and the filling of the rim of the blank shell shown inFIGS. 14A-14C in accordance with the preferred embodiment of theinvention;

FIGS. 16A-16B, 17A-17B, and 18A-18B further illustrate filling of therim of the shell shown in FIGS. 14A-14C and 15A-15C in accordance withthe preferred embodiment of the invention.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

In accordance with one embodiment, the invention provides a porousmetallic scaffold suitable for medical device application, especiallyfor orthopedic implants. In the preferred variant, the porous scaffoldof the invention is a high strength, open cell, metallic foam with poresizes preferably above 100 μm, more preferably, ranging from about 100μm to about 1000 μm, more preferably, ranging from about 300 μm to about500 μm. The porous scaffold is also characterized by high pore volume,ranging from about 50% to about 90%, more preferably, from about 60% toabout 80%. The scaffold has a porous surface. The porous surfaceinterfaces with a bone if the porous scaffold is used in an orthopedicimplant. The porous scaffold is preferably made from biocompatiblemetals, such as titanium, titanium alloys, cobalt-chrome alloy,tantalum, and niobium. The most preferred metals are titanium andtitanium alloys. The preferred titanium alloy is Ti-6Al-4V alloy. Thescaffold may be in a form of a block, a layer, a tissue in-growthsurface or other desirable form or shape.

The porous scaffold P may be attached to a substrate S (FIG. 2).Preferably, the substrate S is a solid metallic substrate. The substrateS and the porous scaffold P are preferably integrated with each other.FIG. 2 shows the scaffold P in the form of a porous layer P₁. It shouldbe understood that the shape of the scaffold P and the substrate S shownin FIG. 2 is purely illustrative, and by no means limiting. Thepreferred thickness d of the porous layer P₁ is ranging from about 0.5mm to about 10 mm, more preferably, from about 1 mm to about 5 mm, yetmore preferably, from about 1 mm to about 2 mm. The thickness of thesolid substrate S may be selected as desired. Preferably, the porouslayer P₁ has a porous surface A with high surface roughness. Anintermediate layer may be present between the substrate S and the porouslayer P₁, for example, for the purposes of bonding the layer P₁ and thesubstrate S.

Preferably, the substrate S and the porous layer P₁ are produced fromthe same metal or alloys of the same metal. The preferred metals arebiocompatible metals, such as titanium, titanium alloys, cobalt-chromealloy, tantalum, and niobium. The most preferred metals are titanium andtitanium alloys. The preferred titanium alloy is Ti-6Al-4V alloy.

The porous scaffold P, with or without the substrate S, is especiallyuseful for medical device applications, such as orthopedic implants. Forexample, the preparation of titanium implants having a porousbone-contacting interface (or tissue in-growth surface) suitable fororthopedic applications presents a number of non-trivial technicalchallenges. In addition to high pore volume, the desirable porous layerfor an orthopedic implant has a rough surface, good pore regularity, andintegrated binding with the underlying solid substrate if such substrateis used. In the past, titanium porous layers that combine these desiredproperties had not been produced.

Thus, in another embodiment, the invention also provides an orthopedicimplant that incorporates the porous scaffold P, which is included inthe implant as a porous bone-contacting surface, porous block, porouslayer or the like. Non-limiting examples of the implants that mayinclude the scaffold P are an acetabular cup implant, vertebral implant,a femoral hip stem implant, femoral and tibial knee joint components,soft tissue attachments, bone defect fillers, shoulder implants,spacers, and any medical device or implant having a surface contacting abone. In addition to the porous scaffold P, the implant may include asolid metallic substrate. Preferably, the solid substrate and the porousscaffold are integrated with each other without cement or any otherexternal binding material. For example, an acetabular cup implant mayinclude a solid shell bearing the scaffold as the bone-contacting poroussurface, block or layer. The porous scaffold facilitates in-growth ofbone tissue into the pores of the scaffold, contributing to a longuseful life of the acetabular cup implant after implantation. Also, theporous surface of the scaffold preferably has high surface roughness,which promotes initial press-fit stability and provides greaterfrictional interference between the porous surface and the bone.

The advantages of the orthopedic implants having the porous scaffold Pare related to the method of forming the scaffold. In anotherembodiment, the invention provides a process for forming porousscaffolds of medical devices, especially orthopedic implants.

FIG. 3 shows a general scheme of the process. First, a desiredbiocompatible metal, such as titanium, is deposited on pyrolyzablepolymer foam by low temperature arc vapor deposition (LTAVD or LTAVdeposition) (Step 1100). LTAVD is a physical vapor deposition (PVD)method that utilizes a high current, low voltage electric arc toevaporate electrically conductive metals. The metal is evaporated inhigh vacuum and is deposited as a thin, highly adherent and densecoating on the desired substrate. The polymer foam includes a polymerweb having an open cell, interconnected structure. The LTAV depositioncreates a thin layer (or skin) of the metal on all surfaces of thepolymer foam. Therefore, the structure of the deposited metal followsthe structure of the polymer web, creating a metal skin over the polymerweb. Controlling various parameters of LTAVD process, especially thetime of the deposition, controls the thickness of the metal skin.

The polymer foam is a low density, high porosity polymer material. Asdescribed above, it serves as a template for the metallic porous layerto be formed. Preferably, the polymer foam is shaped in the same manneras the surface of the desired implant. The polymer foam may be placedaround a solid portion of the future implant before deposition takesplace. More preferably, however, the LTAV deposition is carried out onan unattached piece of the polymer foam. The preferable polymer foamsdecompose with minimal residual contamination upon heating. The foammade from pigmented polyurethane, which does not leave substantialresidue upon decomposition, is preferred.

After the desired thickness of the metal skin is deposited, the polymerfoam coated with the metal skin is heated at temperatures above thedecomposition temperature of the polymer foam in an inert atmosphere(Step 1200). The polymer foam decomposes, leaving behind “green” metalfoam, which is essentially the metal skin formed in the LTAVDdeposition. The term “green”is used to refer to a metal foam that yethas not been strengthened by sintering or other similar techniques.

The next step is pre-sintering of the green metal foam (Step 1300).After pre-sintering, the green foam, which is the weak and thin metalskin, is build up to strengthen the metal foam and to obtain the desiredporosity (Step 1400). The build up involves increasing the thickness ofthe internal surfaces of the pre-sintered foam, which may be termed webthickening. The preferred web thickening method involves applying one ormore layers of metallic powder and binding it to the pre-sintered metalfoam by powder metallurgy techniques. The web thickening may also beaccomplished by LTAV deposition, high temperature PVD or chemical vapordeposition. The web thickening reduces the pore size of the metal foamsince the thickness of the internal pore surfaces increase.

If a single web thickening step provides metal foam with desiredcharacteristics, such as strength and pore size, the foam may besubjected to final sintering. If further web thickening is necessary,the foam is again pre-sintered and the web thickening step is repeated.After the last web thickening step, the metal foam having the desiredthickness, strength, and porosity undergoes a final sintering step,preferably together with the underlying solid metallic substrate (Step1500).

The process will now be described in more detail. The process will bedescribed with reference to the formation of a titanium acetabular cupimplant. However, it should be understood that other biocompatiblemetals, such as titanium alloys, cobalt-chrome alloys, niobium, tantalumand other metals might also be suitable. Likewise, it should beunderstood that similar methods might be used to produce other types ofimplantable medical devices.

To begin manufacturing of the implant, a piece of polyurethane foamhaving a desired thickness, and a shape matching the shape of the futureimplant's bone-interface surface is subjected to LTAV deposition oftitanium. The properties of the polyurethane foam (e.g., porosity,density, and thickness) are important since they may be used to affectthe properties of the final metallic porous layer. Thus, the thicknessof the polyurethane foam determines the thickness of the porous metallayer. In the preferred embodiment, the polyurethane foam has thicknessranging from about 0.5 mm to about 10 mm, more preferably, from about 1mm to about 5 mm, yet more preferably, from about 1 mm to about 2 mm.Also, the porosity of polyurethane foam may be used to control pore sizeand pore volume of the green metal foam and the final porous metallayer. Preferably, the polyurethane foam has pore sizes over 500 μm,more preferably between about 800 μm and about 2000 μm, most preferably,between about 900 μm and about 1100 μm.

As described above, to facilitate bone in-growth, the porous layer ofthe implant preferably has pores size of 100 μm or more. However, if thepore size of the final metal foam is too large, the porous layer maybecome weak because of insufficient structural strength. And, of course,if the pore size is too small, the in-growth of the bone or other tissuemay be retarded. The process of this embodiment of the invention firstproduces the weak green foam with large pore size, and then reduces thepore size in the web thickening step. This methodology allows goodcontrol over the desired pore size and process conditions. The properbalance between strength and in-growth potential is achieved byselecting a combination of porosity of the polyurethane foam and webthickening conditions.

The porosity of the polyurethane foam directly affects the pore size ofthe green metal foam and limits the maximum possible pore size of thefinal porous layer. In an illustrative non-limiting example,polyurethane foam with porosity of 58 pores per cubic inch (ppi) andpore size of 1100 μm may be processed by coating with metal powder toyield final metal foam with pore size of about 600 μm. Under identicalprocessing conditions, the polyurethane foam with porosity of 48 ppi andpore size of 1400 μm yields a final metal foam with the pore size ofabout 900 μm.

Another method to control the pore size of the final foam is to vary thenumber of applications of the titanium powder, which is applied tothicken the green metal foam. The same goal may be accomplished byvarying the particle size of the titanium powder. However, it should beunderstood that if the particle size of the powder is too large, theparticles may not be able to penetrate into the pores of the metal foam.

In another illustrative non-limiting example, 1100 μm polyurethane foammay require two powder layers to produce 600 μm pore size. Increasingthe number of powder layers to three decrease the final metal pore sizeto approximately 400 μm, while applying only one layer of powder wouldresult in final pore size of approximately 800 μm. The thickness and therequired number of layers of the metal powder may be affected by thecharacteristics of the powder particle, such as average size, shape andparticle size distribution.

To a lesser degree, the thickness of the initial LTAVD coating may alsobe varied to affect the pore size of the final metallic foam. In anillustrative non-limiting example, a titanium coating with the thicknessof 25 μm applied by LTAVD to the polyurethane foam with the pore size of1100 μm could contribute to a metal foam with a pore size of 600 μm. Ifthe thickness of the LTAVD coating is increased to 50 μm and all otherprocess parameters are kept the same, the pore size of the final metalfoam would decrease to approximately 550 μm.

FIGS. 4A and 4B show a polyurethane foam shell 110 suitable forproducing a porous layer of an acetabular cup implant. As seen from FIG.4A, the shell 110 matches the shape of a shell of an acetabular cupimplant. The polyurethane foam shell 110 has a first side 111 and asecond side 112. The preferred thickness of the shell 110 is from 1 mmto 2 mm. The shell 110 is subjected to LTAV deposition of titanium. Thepreferred conditions for LTAV deposition of titanium on polyurethanefoam is vacuum of less than 10⁻⁴ torr and electric current setting of130 amperes. LTAVD methodology is described in greater details in U.S.Pat. Nos. 4,351,855, 4,975,230, and 5,011,638, which are incorporatedherein by reference in their entirety. In general, the process of theinvention uses conventional LTAVD methodology. The above-identifiedpatents may be consulted for additional information.

LTAV deposition produces a titanium-coated polyurethane shell 110A (FIG.4C). The deposition creates a titanium skin within the polyurethanefoam. FIG. 5 is a scanning electron microscope (SEM) photograph of thetitanium-coated polyurethane foam at 25× and 1000× magnifications. Thetitanium skin coats internal and external surfaces of the polymer web.Preferably, the thickness of the titanium skin is from about 1 μm toabout 50 μm. More preferably, the thickness of the titanium skin isranging from about 10 μm to about 35 μm. Most preferably, the thicknessof the titanium skin is approximately 10 Am for the first side 111, andabout 35 μm for the second side 112 of the shell 110. In the LTAVDprocess, the thickness of the titanium skin is varied by turning overthe polyurethane foam shell and coating the second side 112 for longerperiod of time.

A blank metal shell 210 of an acetabular cup implant serves as asubstrate for final metal foam (FIG. 6). The blank shell 210 is madefrom solid titanium. The blank titanium shell 210 has a top surface 211and a bottom surface 212 (FIG. 6). After the LTAV deposition, thetitanium-coated polyurethane foam shell 110A is wrapped around the blankshell 210 with the second side 112 of the coated polyurethane foam shell110 facing the top surface 211 of the blank titanium shell 210 (FIG. 7).Then, an assemblage 230 of the blank shell 210 and the attached shell110A is heated to decompose polyurethane.

The preferred heating set up is shown in FIG. 8. As seen from FIG. 8,the assemblage 230 is placed in a retort 310 equipped with an argoninlet 312, a gas exhaust 314, and a thermocouple 316, purged with argon,and transferred to a furnace 330. The furnace 330 is equipped withheating elements 331. Inside the furnace 330, the assemblage 230 ismaintained under the argon atmosphere to prevent oxidation of titanium.

The furnace 330 is maintained at a temperature substantially above thedecomposition temperature of polyurethane (177° C.). The preferredfurnace temperature is from about 1050° C. to about 1150° C., the morepreferred furnace temperature is from about 1055° C. to about 1075° C.Because of the high temperature in the furnace, the assemblage 230 israpidly heated, decomposing polyurethane in the shell

The decomposition of polyurethane results in a build-up of decompositiongases inside the titanium skin of the shell 110A. Referring to FIG. 7,the decomposition gases rupture the titanium skin on the first side 111of the shell 110A, creating cracks to allow gases to escape. Theinventors found that the thickness of titanium skin of the side 111 isimportant to control the escape of the decomposition gases. Theinventors also found that it is important to rapidly heat the assemblage230. Rapid heating is believed to contribute to minimizing residue.

Once the temperature inside the retort 310 exceeds the decompositiontemperature of polyurethane by about 400° C., the burn-off cycle iscomplete. The complete polyurethane burn-off takes approximately 5 to 10minutes. The retort 310 is removed from the furnace 330, and theassemblage 230 is allowed to cool to room temperature in an argonatmosphere.

The burning off of polyurethane produces green titanium foam 110B on thesurface of the shell 210 (FIG. 9). FIG. 10A shows a SEM photograph ofthe green titanium foam after removal of the polyurethane foam. Thegreen titanium foam typically has pores similar to or slightly largerthan the pores of the starting polyurethane foam. Typically, the greentitanium foam has pore sizes 1% to 5% greater than the pore sizes of thepolyurethane foam, often about 3% greater. For example, if the pore sizeof the polyurethane foam of the shell 110 is 1000 μm, the pore size ofthe green titanium foam 110B may be 1000-1050 μm.

The green titanium foam 110B is fragile. To strengthen the green foambefore it can be thickened to the desired pore size/pore volume, theassemblage is pre-sintered in a conventional lab furnace. The preferredpre-sintering temperature is from about 1260° C. to about 1370° C., morepreferably, it is about 1315° C. After pre-sintering, the green foam maynow be used for further processing. FIG. 10B shows a SEM photograph ofthe pre-sintered titanium foam.

The next step in the process is to thicken and strengthen the titaniumfoam and to achieve a desired pore volume and pore size. The morepreferred pore size for the porous layer of an implant is from about 300μm to about 500 μm. In one variant, LTAVD process may be used to extendthe titanium web. Preferably, the titanium web is extended through apowder metallurgy process. Powder metallurgy involves binding metalparticles into a solid whole and/or applying a metal powder to asurface, usually a metallic surface, and bonding the powder to thesurface by heating.

A scheme of the preferred powder metallurgy process is illustrated inFIG. 11. The pre-sintered metal foam is sprayed with a solution of abinder (Step 1310). The binder is used to provide a temporary bondbetween the surfaces of the pre-sintered foam and external titaniumpowder. Preferably, an atomized (ultra fine) binder, preferably in theform of a mist, is delivered to the foam by an ultrasonic atomizingnozzles system. A layer of binder forms on all internal and externalsurfaces of the pre-sintered foam. In the most preferred embodiment, thenozzle employs a high frequency (e.g., 65 KHz) sound wave to atomize thesolution of the binder into droplets with an average size of from about20 μm to about 80 μm, more preferably, from about 30 μm to about 40 μm,and to deliver the droplets to the foam at a velocity of from about 0.6to about 1.2 fps. Because of the small size of the binder droplets, thebinder reaches substantially every surface inside and outside thetitanium foam. Also, the use of the ultrasonically-produced ultra finebinder allows delivery of the binder inside the foam without bridgingthe pores.

Any binder suitable for orthopedic applications, such as fish glue andthe like may be used. The preferred binder is a 2% aqueous solution ofmethyl cellulose with a viscosity of approximately 25 cps. Methylcellulose leaves less carbon residue on the titanium foam than a fishglue after the binder is decomposed in sintering.

After the binder is sprayed, a powder of titanium particles is sprayedon the foam covered with the binder (Step 1320). It is desired todeliver the powder to every surface of the foam. For this reason, thesize of the titanium particles is smaller than the pore size of themetal foam so that the particles may reach inside the foam withoutbridging the pores. The preferred titanium powder has a particle size offrom about 20 μm to about 100 μm, more preferably from about 40 μm toabout 80 μm. A powder spray delivery system is used to increase theparticle momentum so that the particles may get into the bottom layer ofthe pre-sintered foam. As the titanium powder comes in contact with thefoam, the binder ties the powder to the surfaces of the foam. After thepowder is applied, the excess of the powder is removed by air spraying(Step 1330), and the metal foam is sintered (Step 1340), producingthickened metal foam with pores smaller than the pores of the metal foambefore web thickening. If the desired pore size is achieved with asingle application of titanium powder, the sintering may be final andthe thickened foam is the final metal foam. Alternatively, the foam issubjected to another pre-sintering, and the binder spraying/powderapplication is repeated until the desired pore size is obtained.

The final sintering is used to improve the strength of the porous layerand the bond between the porous layer and the underlying solidsubstrate. After the final sintering, the metal foam is integrated withthe underlying substrate into a unitary component. Preferably, the finalsintering is carried out under high vacuum with 9° C. per minute ramprate. The preferred temperature for final sintering is from about 1425°C. to about 1530° C., more preferably, the final sintering is done atapproximately 1500° C.

The web thickening converts the foam 110B into a thickened foam 110C(FIG. 12). As described above, the process of the invention may includeone, two or more web thickening steps. Thus, the thickened foam 110C maybe the final metal foam or an intermediate foam. FIG. 13 shows a SEMphotograph of the final titanium foam (porous layer for an implant)shown at 25× magnification. The final porous titanium foam preferablyhas the pore volume from about 50% to about 90%, more preferably from 60to 80%, and the pore size from about 100 μm to about 1000 μm, morepreferably, from about 300 μm to about 500 μm. If desired, the finalmetal foam may be coated with a biocompatible coating.

In other embodiments, the invention also provides an acetabular cupimplant that incorporates a porous metallic layer and a preferredstructure of such implants, as well as a method of making suchacetabular cup implant that improves adhesion and stability of theporous metallic layer.

In the preferred embodiment, the metal foam 110C is not immediatelysubjected to final sintering. Instead further processing is carried outto improve the stability of the metal foam. Referring back to FIG. 6,the blank titanium shell 210 may have a rim 215. FIGS. 14A-14C showadditional details of the structure of the rim 215. The rim 215 includesa ledge 217 and a wall 218 (FIGS. 14A and 14B). The wall 218 has a topsurface 218 a (FIG. 14C). The ledge 217, the wall 218 and a section ofthe shell 210 define a recess 216 (FIG. 14C). After the blank shell 210is processed as described above, the porous foam 110C and the shell 210are integrated into a unitary component 230A having the rim 215 (FIGS.15A and 15B). When an acetabular cup implant having the foam 110C isimplanted, significant forces are exerted upon the foam, for example asshown by the arrow A (FIG. 15B). If the foam 110C is not supported frombelow, the application of such forces may lead to disintegration of thefoam, for example as shown by the arrow B. Therefore, in the preferredembodiment, the component 230A is further processed to improve stabilityof the porous metallic foam 110C. It should be understood that similarmethods might be used for other implantable medical devices.

Referring to FIG. 15C, the foam 110C has sections 110.1 and 110.2. Thesection 110.2 extends into the recess 216 below the top surface 218 a ofthe wall 218. The component 210A is sprayed with a binder, and the rim215 is immersed into titanium powder approximately at the level of thetop surface 218 a of the wall 218. The immersion causes the powder tofill the recess 216. The particle size of the powder is selected so thatthe powder may enter into the pores of the foam 110C. Most of thesection 110.1 of the foam 110C is not immersed in the powder. However,the section 110.2 is below the powder fill level (top surface 118 a).Thus, the powder fills both the recess 216 and the pores of the section110.2. A vibrational treatment may be used to facilitate the filing ofthe pores in the section 100.2 of the foam 110C.

Once the rim 215 is full, the component 230A is subjected to finalsintering, producing a component 230B having a filled rim 215 a (FIGS.16A and 16B). In one variant, the filling of the rim 215 may be combinedwith additional web thickening step. If so, upon final sintering, thefoam 110C is converted to a thicker final titanium foam 110D (FIG. 16A).In another variant, the foam 110C may be the final titanium foam.

As seen in FIG. 16B, in the component 230B, the prior opening 216 andthe section 110.2 are converted to a substantially solid metal block 216a, while the section 110.1 of the final foam 110C or 110D remainsporous. The block 216 a is integral with the surface 211 of the startingblank shell 210, the ledge 217, and the wall 218. The block 216 asupports the metal foam 110C or 110D, providing improved stability forthe metal foam.

Generally, the component 230B, especially the filled rim 215 a, may bemachined as desired to obtain a shell of an acetabular cup implanthaving a desired shape and/or dimensions. For example, after finalsintering, the component 230B may be machined as shown by arrows C inFIG. 16A. The machining removes portions of the rim 215 and the block216 a, and produces a component 230C shown in FIGS. 17A-17B. As seen inFIG. 17B, the block 216 a is converted into a substantially solid layer216 b, which lies above a ledge portion 217 a. After machining, thelayer 216 b supports the metal foam 110.1 (FIG. 17B). The ledge portion217 a may be removed by further machining, for example as shown byarrows D in FIG. 17A. In the resulting machined component 230D (FIGS.18A and 18B), the final metal foam is supported by the substantiallysolid layer 216 b.

EXAMPLE 1 Preparation of an Acetabular Cup Implant

It should be understood that while an acetabular cup implant isillustrated, this should not be considered a limitation on the scope ofthe invention.

A. Formation of Green Foam.

A block of polyurethane (PU) foam (Foamex, 950 μm, 58±2 ppi) is machinedto a thickness of 1.5 mm and desired size and shape matching a shell ofan acetabular cup. The resulting PU foam shell is slightly oversize(˜3%) with respect to the size of the shell of the acetabular cup. ThePU foam shell is subjected to LTAV deposition of titanium at 93° C. Theside of the PU foam shell that will face the acetabular cup shell (IDside) is subjected to deposition for approximately 53 hours, whereas thedeposition of titanium on the other side of the PU foam shell (OD side)is carried out for approximately 15 hours. The deposition is concludedwhen the thickness of the titanium coating reaches approximately 35 μmon the ID side and approximately 10 μm on the OD side of the PU foamshell. After the deposition is complete, the titanium-coated PU foamshell is attached to a blank of acetabular cup shell with the ID side ofthe PU foam shell facing the surface of the blank shell. The blank shellis made of titanium. The assembly of the blank titanium shell and thetitanium-coated PU foam shell is placed in a retort equipped with anargon inlet and a thermometer. The retort and the assembly are placedinto a furnace maintained at 1071-1121° C. under a continuous flow ofargon at 40 ft³/hour. After 5 to 10 minutes, the temperature in theretort reaches 550-600° C., indicating the complete burn-off ofpolyurethane. The retort is removed from the furnace and cooled to roomtemperature. The flow of argon through the retort is maintained duringthe cooling to minimize oxidation of the green foam. The resulting greenfoam on the surface of the shell has a pore size of approximately980-1000 μm.

B. Pre-Sintering of the Green Foam.

To pre-sinter the green titanium foam, the shell is placed in a vacuumoven and the air is evacuated from the oven. Once the vacuum reaches10⁻⁵ torr, the oven is heated to 427° C. at a ramp rate of approximately8.3° C. per minute. The oven temperature is maintained at 427° C. forapproximately 15 minutes. The heating is resumed at the same ramp rateuntil the temperature reaches 1316±22 ° C. The shell is maintained inthe oven at 1316±22° C. for approximately 2 hours to complete thepre-sintering. The oven is cooled to room temperature, and the shellwith the pre-sintered foam is removed.

C. Thickening of the Foam.

The pre-sintered titanium foam is sprayed with an ultra fine mist of abinder (2% aqueous solution of methyl cellulose, 25 cps, droplet size30-40 μm). The stream of the binder is delivered by a Sonotek ultrasonicsprayer nozzle. After spraying, the binder is distributed throughout thefoam. The binder-covered foam is then treated with titanium powder (TiCP2, 40-80 μm particle diameter). The powder is sprayed by a metalpowder sprayer onto and into the foam covered with methyl cellulosesolution. The foam is air dried, and the shell with the foam is againpre-sintered as described above. After second pre-sintering, the cup isagain treated with the binder solution, and another application oftitanium powder is sprayed into the foam. The shell is then transferredto a vibratory table having a container filled with titanium powder(CP2, ˜45 μm particle diameter). The vibratory table is turned on. Thecup is immersed into the container until a rim of the shell is filledwith the powder. The cup is removed from the container and placed on thevibratory table for 5 minutes. If the rim is not full after thevibrational treatment, the shell is again immersed into the titaniumpowder and the vibrational treatment is repeated for another 2 minutes.After the rim is full, the cup is air-dried for 12 hours at roomtemperature. The shell is ready for final sintering.

D. Final Sintering.

The shell is placed into a vacuum oven. Once the vacuum reaches 10⁻⁵torr, the oven is heated at a ramp rate of 8.3° C. per minute to 427° C.The shell is kept at 427° C. for approximately 15 minutes. The heatingis resumed at the same ramp rate until the temperature reaches 1316±22°C., and the shell is maintained at this temperature for approximately 2hours. The oven temperature is raised to 1496±9° C. within 10 minutes.The shell is then maintained at this temperature for approximately 90minutes to complete the sintering process. The shell is cooled to roomtemperature under vacuum, and then the oven is filled with inert gas.The pore size of the final foam is approximately 500-520 μm. The cup ismachined to desired specifications.

EXAMPLE 2

The shell is processed as in the Example 1, but instead of the twopowder treatments, the foam is thickened by vapor deposition (e.g.,LTAVD, PVD, or CVD) of titanium until a 150-200 μm layer is deposited.

EXAMPLE 3

The shell is processed as in the Example 1, but the pore size of thepolyurethane foam is 924 μm±89 μm, the pore size of the green foam afterLTAV deposition is 967 μm±82 μm, and the pore size of the final foam is614 μm±67 μm.

EXAMPLE 4

The shell is processed as in the Example 1, but only one powdertreatment step is carried out, the pore size of the polyurethane foam(Crest, S-50 natural color) is 600 μm±50 μm, the pore size of the greenfoam after LTAV deposition is 630 μm±45 μm, and the pore size of thefinal foam is 480 μm±42 μm.

EXAMPLE 5

Protocol for Measuring Pore Sizes.

A sample of foam having a thickness of 1 to 2 mm is imaged by a ScanningElectron Microscope producing a field of view depicting approximately 5mm×5 mm of the coated surface. Measurements are taken throughout tendistinct fields of view from a single sample. The distance between theweb surfaces surrounding a pore is measured on all complete pores withina field of view. A complete pore is defined as one having allsurrounding webs and pores intact. There are usually 2 to 4 completepores within a single field of view, typically yielding an overallsample size of 20 to 40 readings. The mean and sample standard deviationare calculated and reported in microns.

Unless stated to the contrary, any use of the words such as “including,”“containing,” “comprising,” “having” and the like, means “includingwithout limitation” and shall not be construed to limit any generalstatement that it follows to the specific or similar items or mattersimmediately following it. Also, if a range is described in thespecification and/or recited in the claims, the description/recitationof the range covers every data points within the range, as well as thebeginning and ending points of the range. Each such data point, as wellas the range defined thereby, should be considered as separatelydisclosed and/or claimed.

Although the invention herein has been described with reference toparticular embodiments, it is to be understood that these embodimentsare merely illustrative of the principles and applications of thepresent invention. It is therefore to be understood that numerousmodifications may be made to the illustrative embodiments and that otherarrangements may be devised without departing from the spirit and scopeof the present invention as defined by the appended claims.

1. A porous metal scaffold for use in an implantable medical devicecomprising: a porous biocompatible metal network having an open cellstructure wherein the openings of each cell are defined by webs with ametal skin surrounding an empty core, the metal skin having a porouslayer of biocompatible metal particles, the metal particles are bondedto the metal skin and other particles, the particle covered websdefining each cell having a pore size of 100 to 1000 μm for tissueingrowth, the porous layer of metal particles including portions havinga diameter of 40 to 80 microns.
 2. The porous metal scaffold as setforth in claim 1 wherein the metal skin surrounding the empty core hasopenings therein.
 3. The porous metal scaffold as set forth in claim 1wherein the metal of the metal skin and metal particles are selectedfrom the group consisting of titanium, titanium alloy, cobalt chromealloy, niobium and tantalum.
 4. The porous metal scaffold as set forthin claim 1 wherein the final pore volume of the metal scaffold is 50% to90%.
 5. The porous metal scaffold as set forth in claim 1 wherein theporous metal network has a thickness of 0.5 mm to 5 mm.
 6. The porousmetal scaffold as set forth in claim 1 wherein the webs form acontinuous metal skeleton of the porous metal scaffold.
 7. The porousmetal scaffold as set forth in claim 1 wherein the porous layer of metalparticles has neck portions extending between adjacent particles.
 8. Aporous metal scaffold for use in an implantable medical devicecomprising: a porous biocompatible metal network having an open cellstructure wherein the openings of each cell are defined by webs with ametal skin surrounding an empty core, the metal skin having at least oneporous layer of biocompatible metal particles, the metal particles arebonded to the metal skin and other particles, the porous metal networkdefined by the metal skin covered by the porous layer of metal particlesexhibit a pore size of 100 to 1000 μm for tissue ingrowth, each cellhaving a profile formed by surfaces of the porous layer of bondedparticles having portions between 40 to 80 μm in diameter.
 9. The porousmetal scaffold as set forth in claim 8 wherein the metal skinsurrounding the empty core has openings therein.
 10. The porous metalscaffold as set forth in claim 8 wherein the metal of the metal skin andmetal particles are selected from the group consisting of titanium,titanium alloy, cobalt chrome alloy, niobium and tantalum.
 11. Theporous metal scaffold as set forth in claim 8 wherein the final porevolume of the metal scaffold is 50% to 90%.
 12. The porous metalscaffold as set forth in claim 8 wherein the porous metal network has athickness of 0.5 mm to 5 mm.
 13. The porous metal scaffold as set forthin claim 8 wherein the webs form a continuous metal skeleton of theporous metal scaffold.
 14. The porous metal scaffold as set forth inclaim 8 wherein the web surface covered by particles has neck portionsextending between adjacent particles.